Injectables and resurfacing techniques: Lasers in aesthetic surgery

Historical perspective

Lasers have been used in medicine for several decades. Alexander et al . noted that Albert Einstein was the first to describe the concept behind the laser in the theory of stimulated emission of radiation in 1917. Theodore Maiman developed the first laser light with use of a ruby crystal in 1960 and shortly thereafter, lasers were being used in ophthalmology. During the next several years, practitioners in many other medical fields, including dermatology, otolaryngology, neurosurgery, and gynecology, incorporated lasers into their practices. Selective photothermolysis is the theoretical foundation that allowed the development of lasers designed to target specific tissues. The first of these lasers was the pulsed dye laser, with a wavelength of 577 nm, one of the absorption peaks of hemoglobin. This device was applied primarily for vascular lesions. Leon Goldman, a dermatologist considered by many to be the father of laser medicine, pioneered much of the early work with cutaneous lasers and published his experience in the treatment of vascular birthmarks with the ruby laser.

The CO ₂ laser, developed in 1964, was initially a continuous wave laser. It was designed with a highly focused handpiece to cut skin. Its injury was associated with a bloodless field. It also appeared to generate less edema, possibly because the laser sealed lymphatics. Sterling Baker, an ophthalmologist, is credited with having performed the first laser blepharoplasty in 1984. The CO ₂ laser was also used as an ablative tool for cutaneous lesions, but early lasers had limited control of energy parameters, leading to frequent thermal injury and scarring.

Changing the delivery of laser energy from a continuous to a pulsed mode was key in the early improvement of the resurfacing laser. Electronic shutters were developed to interrupt the continuous wave and produce intermittent bursts of laser energy (known as pulses), thereby decreasing the laser exposure time. The original pulses were 0.1–1 s in duration, still too long an exposure time to avoid scar formation unless very low powers were applied. Superpulse lasers were then developed, which delivered shorter pulse durations (pulse widths) and higher power, but still led to an unacceptable incidence of complications. Ultrapulse technology provided an increase in power seven times that of the superpulse lasers, and the pulse width was decreased to less than 1 ms. These early advances helped limit the thermal damage and associated scar formation, allowing the CO ₂ laser to be used effectively as a resurfacing tool. In 1991, Fitzpatrick et al . used this laser to treat actinic damage of the face and actinic cheilitis and also noted improvement in wrinkles. These observations led to the use of laser resurfacing with short-pulsed, high-energy CO ₂ to treat photodamaged skin and acne scars.

The introduction of short-duration Er:YAG lasers in the mid-1990s offered another option for resurfacing, either alone or in combination with the CO ₂ laser. Laser resurfacing grew in popularity and gained widespread acceptance among aesthetic plastic surgeons. As more problems with scarring and hypopigmentation were reported, the early enthusiasm toward ablative CO ₂ waned.

In an effort to expand treatment to all skin types and to improve safety, while reducing the frequency of scarring and loss of pigmentation, fractional laser resurfacing was introduced by Reliant Technologies in 2003 as a new class of therapy. Fractional photothermolysis has become an important modality in management of a number of skin conditions and photoaging. Unlike full-surface flat-beam resurfacing, fractional resurfacing damages specific microtreatment zones within the target area. It seems that almost any laser wavelength can be fractionated, and the initial non-ablative fractional lasers have been shown to be safe and effective for treatment of fine wrinkles, pigmentation, and acne scars.

Deeper rhytids and more extensive solar elastosis have not responded as well to non-ablative wavelengths and, hence, fractional ablative CO ₂ and Er:YAG lasers have been developed that retain much of the safety of non-ablative fractional resurfacing but push the efficacy closer to non-fractional laser ablation. Since the introduction of the first fractional laser in 2003, numerous devices delivering fractionated treatment with a variety of different capabilities have appeared on the market. More importantly, in the aesthetic arena, where safety is of paramount importance, fractional laser resurfacing has resonated widely with both patients and physicians alike and has become a leading trend in minimally invasive cosmetic surgery.

Inflammation

Tissue injury leads to parenchymal cell damage and the extravasation of blood constituents. The blood clot that forms re-establishes hemostasis and provides a provisional matrix for cell migration. Vasoactive and inflammatory mediators are then generated by platelets (platelet-derived growth factor; PDGF), the coagulation cascade, the activated complement pathway (C3a and C5a), and injured or activated parenchymal cells. Vasoconstriction begins within seconds of the injury and lasts 10–15 min and occurs as a result of epinephrine being released into the peripheral circulation, stimulation of the sympathetic nervous system, and release of norepinephrine. Vasodilatation and capillary leak are mediated by a variety of factors including leukotrienes, prostaglandins, kinins, histamine, and complement factors C3a and C5a. Leukocyte migration is stimulated by components of the extracellular matrix and several inflammatory mediators. Neutrophils cleanse the wound of foreign particles and bacteria and are then extruded with the eschar or phagocytosed by macrophages. In response to specific chemoattractants, monocytes infiltrate the wound and become activated macrophages that contribute to the coordination of wound healing. Specifically, angiogenesis, fibroblast migration and proliferation, collagen production, and wound contraction are all directed by macrophages.

The wound healing process is regulated, in a large part, by the ordered production of cytokines that control gene activation responsible for cellular migration and proliferation and synthetic activities.

Proliferation

This phase includes epithelialization, angiogenesis, granulation tissue formation, and collagen deposition.

Epithelialization was discussed above in the previous section. Angiogenesis, stimulated by tumor necrosis factor alpha (TNF-α), is marked by endothelial cell migration and capillary formation. The migration of capillaries into the wound bed is critical for proper wound healing. Much of the angiogenesis occurs during the early phase of wound healing and involves the sprouting of endothelial cells from post-capillary venules. The granulation phase and tissue deposition require nutrients supplied by the capillaries, and failure of this to occur results in a chronically unhealed wound. The final part of the proliferative phase is granulation tissue formation. Fibroblasts migrate into the wound site from the surrounding tissue, become activated, and begin synthesizing collagen and proliferating. PDGF and EGF are the main signals to fibroblasts and are derived from platelets and macrophages. PDGF expression by fibroblasts is amplified by autocrine and paracrine signaling. Fibroblasts already located in the wound site (termed “wound fibroblasts”) will begin synthesizing collagen and transform into myofibroblasts for wound contraction (induced by macrophage-secreted TGF-β1); they have less proliferation compared with the fibroblasts coming in from the wound periphery. In response to PDGF, fibroblasts begin synthesizing a provisional matrix composed of collagen type III, glycosaminoglycans, and fibronectin.

Remodeling

This phase comprises both wound contraction and collagen remodeling.

Wound contraction is the result of specialized fibroblasts that express α-actin (myofibroblasts) and their interaction with cytokines and the extracellular matrix (ECM). During the second week of healing, fibroblasts assume a myofibroblast phenotype characterized by large bundles of actin-containing microfilaments along the cytoplasmic surface of the cell membrane. Fibroblasts maintain contact with collagen matrix by integrin receptors. Contraction requires stimulation by TGF-β, PDGF, and myofibroblast–ECM interaction via integrin receptors. As myofibroblasts contract, they exert force on the ECM and subsequently the wound margin. Scar remodeling predominates after ≈21 days. There is no net increase in collagen content despite an increase in tensile strength. Collagen production continues, although at a slower rate, and there is an equal rate of collagen breakdown by collagenases. As the wound matures, disorganized fine collagen fibers are replaced with thicker fibers arranged parallel to skin stresses. The percentage of type III collagen gradually decreases, as does the quantity of water and proteoglycans. The duration of the maturation process varies depending on how long the wound remains open.

The edge of migrating epithelium marks transition between inflammation and fibroproliferation. In the center, where the wound is open, chronic bacterial invasion provides persistent stimulus for inflammation. This tissue contains inflammatory cells, a high concentration of immature vessels, and the components of a provisional matrix (collagen type I, fibrin, and fibronectin). When the inflammatory response is prolonged, this tissue looks like “granulation tissue”. Once epithelium covers the wound, the inflammatory stimulus is eliminated, and fibroblasts predominate. Further behind the migrating epithelium, there are fewer fibroblasts, indicating a more mature wound. Epithelial cells appear to be the source of a stimulus for inflammatory cells to undergo apoptosis. However, if the inflammatory phase continues longer than 2–3 weeks, then this stimulus may be lost, and hypertrophic scarring may result.

Molecular basis of light–tissue interaction

In any light–tissue interaction (LTI), the thermal or photochemical effects depend on the local absorbed energy density at the target. Spatial localization of temperature elevation is possible when (1) the absorption coefficient of the target exceeds that of surrounding tissue (selective photothermolysis); (2) when the “innocent bystander” tissues are cooled so that their peak temperatures do not exceed some damage threshold; or (3) by applying very small (usually <500 μm in diameter) beamlets or microbeams (i.e., fractional methods). Localized heating, for example, in telangiectases and lentigines, follows from the relative excess of Hgb and melanin, respectively, in the lesions versus surrounding skin. In contrast, “non-fractional” mid-infrared (MIR) lasers spatially confine temperature elevation by using heating and cooling schemes that allow for selective dermal heating.

Targeting discrete chromophores offers advantages over targeting tissue water, especially where the ratio of light absorption between the chromophore and surrounding tissue is large (i.e., >10). For example, at least in lighter-skinned patients, targeting dermal Hgb can be achieved with minimal surface cooling. Whereas cooling is desirable, even in these cases, the primary indication is analgesia rather than epidermal protection. Also, the risk of a severe injury to the skin is lessened, as there is no bulk heating. Lastly, because temperature elevations are localized, there can be less pain than with devices targeting ubiquitous tissue water. Thermal injury is determined by time/temperature combinations. Protein denaturation is dependent linearly on exposure time and exponentially on temperature: that is, cell death is more sensitive to temperature than time. Most devices for rejuvenation are based on photothermal or “electrothermal” mechanisms: that is, the conversion of light or electrical energy to heat.

Two processes govern all interactions of light with matter: absorption and scattering. The absorption spectra of major skin chromophores dominate laser–tissue interactions. If tissues were clear, then only absorption would be required to characterize light propagation in skin. However, the dermis appears white because of light scatter (milk is a reasonable model for the dermis in terms of scattering). Scattering is responsible for much of light’s behavior in the skin (beam dispersion, spot size effects, etc.). The main scattering wavelengths are between 400 and 1200 nm (i.e., those where tissue water absorption is poor).

There are three chromophores of interest in skin (water, blood, and melanin). Water makes up about 65% of the dermis and lower epidermis. There is some water absorption in the ultraviolet (UV) spectrum. Between 400 and 800 nm, water absorption is quite small (which is consistent with our real-world experience that visible light propagates quite readily through a glass of water). Beyond 800 nm, there is a small peak at 980 nm, followed by larger peaks at 1480 and 10,600 nm (CO ₂ ). The maximal absorption for water is at 2940 nm (Er : YAG). ( Fig. 8.4.2 ).

Reaction types